Hydrolysable hydrogels for controlled release

ABSTRACT

The present invention relates to a biodegradable hydrogel comprising bonds which are hydrolysable under physiological conditions. More particularly, the hydrogel consists of two interpenetrating polymer networks interconnected to one another through hydrolysable spacers. In addition, the invention relates to a method for the preparation of a hydrogel, wherein macromolecules, e.g. polymers, which contain bonds which are hydrolysable under physiological conditions, are cross-linked in an aqueous solution.

[0001] The present invention relates to hydrogels which have a goodcontrolled release behaviour, and to processes to prepare suchhydrogels.

[0002] The fast developments in the field of molecular biology andbiotechnology have made it possible to produce a large number ofpharmaceutically interesting products in large quantities. For instance,pharmaceutically active peptides and proteins can suitably be used asdrugs in the treatment of life-threatening diseases, e.g. cancer, and ofseveral types of viral, bacterial and parasital diseases; in thetreatment of e.g. diabetes; in vaccines, e.g. for prophylactic aims, andfor anticonception purposes. Especially the specialized biologicalactivities of these types of drugs provide tremendous advantages overother types of pharmaceutics.

[0003] To illustrate the fast developments, it has been reported (seee.g. Soeterboek and Verheggen, Pharm. Weekblad 130 (1995) 670-675) thatin the United States of America, about 275 biotechnological products arein phase IV studies, while more than 500 products are underinvestigation.

[0004] Examples of (recombinant) proteins, which are considered veryinteresting from a pharmacological point of view, are cytokines, such asinterleukines, interferons, tumor necrosis factor (TNF), insulin,proteins for use in vaccines, and growth hormones.

[0005] Due to their nature, proteins and proteinaceous products,including peptides, which group of products will be referred to asprotein drugs herein-below, cannot be administered orally. Theseproducts tend to degrade rapidly in the gastro-intestinal tract, inparticular because of the acidic environment and the presence ofproteolytic enzymes therein.

[0006] Moreover, to a high extent protein drugs are not able to passendothelial and epithelial barriers, due to their size and, generally,polar character.

[0007] For these reasons, protein drugs have to be brought in the systemparenterally, i.e. by injection. The pharmacokinetical profile of theseproducts is, however, such that injection of the product per se requiresa frequent administration. For, it is a known fact that proteinaceousmaterial is eliminated from the blood circulation within minutes.

[0008] In other words, since protein drugs are chemically and/orphysically unstable and generally have a short half-time in the human oranimal body, multiple daily injections or continuous infusions arerequired for the protein drug to have a desired therapeutic effect. Itwill be evident that this is inconvenient for patients requiring theseprotein drugs. Furthermore, this type of application often requireshospitabilization and has logistic drawbacks.

[0009] In addition, it appears that at least for certain classes ofpharmaceutical proteins, such as cytokines which are presently used ine.g. cancer treatments, the therapeutic efficacy is strongly dependenton effective delivery, e.g. intra- or peritumoral. In such cases, theprotein drugs should be directed to the sites where their activity isneeded during a prolonged period of time.

[0010] Hence, there is a need for delivery systems which have thecapacity for controlled release. In the art, delivery systems consistingof polymeric networks in which the proteins are loaded and from whichthey are gradually released have been proposed.

[0011] More in detail, at present, two major types of polymeric deliverysystems can be distinguished: biodegradable polymers andnon-biodegradable hydrogels.

[0012] Biodegradable polymers, e.g. polylactic acid (PLA) and copolymersof PLA with glycolic acid (PLGA), are frequently used as deliverysystems for proteins.

[0013] Proteins can be incorporated in pharmaceutical delivery systems,e.g. microspheres, by a variety of processes. In vitro and in vivo,usually a biphasic release profile is observed: an initial burstfollowed by a more gradual release. The burst is caused by proteinaceousmaterial present at or near the surface of the microspheres and byproteinaceous material present in pores. The gradual release is ascribedto a combination of diffusion of the proteinaceous material through thematrix and degradation of the matrix. Especially for larger proteinsdiffusion in these matrices is negligible, so that the release dependson the degradation of the polymer. The degradation can be influenced bythe (co)polymer composition. A well-known strategy to increase thedegradation rate of PLA is co-polymerization with glycolic acid.

[0014] Although delivery systems based on biodegradable polymers areinteresting, it is very difficult to control the release of theincorporated protein. This hampers the applicability of these systems,especially for proteins with a narrow therapeutic window, such ascytokines and hormones. Furthermore, organic solvents have to be usedfor the encapsulation of the protein in these polymeric systems.Exposure of proteins to organic solvents generally leads todenaturation, which will affect the biological activity of the protein.Furthermore, the very stringent requirements of registration authoritieswith respect to possible traces of harmful substances may prohibit theuse of such formulations of therapeutic drugs in human patients.

[0015] Also hydrogels are frequently used as delivery systems forproteins and peptides. Hydrogels can be obtained by crosslinking awater-soluble polymer yielding a three-dimensional network which cancontain large amounts of water. Proteins can be loaded into the gel byadding the protein to the polymer before the crosslinking reaction iscarried out or by soaking a preformed hydrogel in a protein solution.So, no (aggressive) organic solvents have to be used to load thehydrogels with protein molecules.

[0016] In contrast with the biodegradable polymers, the release ofproteins from hydrogels can be easily controlled and manipulated byvarying the hydrogel characteristics, such as the water content and thecrosslink density of the gel. However, a major disadvantage of thecurrently used hydrogel delivery systems is that they are notbiodegradable. This necessitates surgical removal of the gel from thepatient after the release of the protein in order to preventcomplications of inclusion of the empty hydrogel material (wound tissueis frequently formed).

[0017] Biodegradable hydrogels have been used in the preparation ofdelivery systems for protein drugs. One of these systems comprisescrosslinked dextrans obtained by coupling glycidyl methacrylate (GMA) todextran, followed by radical polymerization of an aqueous solution ofGMA-derivatized dextran (dex-GMA). In this respect, reference is made toVan Dijk-Wolthuis et al. in Macromolecules 28, (1995), 6317-6322 and toDe Smedt et al. in Macromolecules 28, (1995) 5082-5088.

[0018] Proteins can be encapsulated in the hydrogels by adding proteinsto a solution of GMA-derivatized dextran prior to the crosslinkingreaction. It appeared that the release of the proteins out of thesehydrogels depends on and can be controlled by the degree of crosslinkingand the water content of the gel (Hennink et al., J. Contr. Rel. 39,(1996), 47-57).

[0019] Although the described crosslinked dextran hydrogels wereexpected to be biodegradable, these hydrogels are rather stable underphysiological conditions. This can is further elaborated in Example 5.It is shown among other that the dissolution time of dextran hydrogelsobtained by polymerization of dextran derivatized with glycidylmethacrylate (DS=4) had a dissolution time of about 100 days. Dextranhydrogels, wherein the dextrans have a higher degree of substitution,did not show any signs of degradation during 70 days, even at extremeconditions.

[0020] The object of the present invention is to provide a slow orcontrolled release delivery system which does not possess theabove-mentioned disadvantages, and especially does not require the useof organic solvents, does not show the undesired and uncontrollableburst effects, and do not possess a poorly controllable releasebehaviour. The present invention aims to combine the advantages of bothtypes of known delivery systems, viz. a system, (bio)degradable underphysiological conditions— either chemically and/or enzymatically—, withcontrolled protein drug release.

[0021] The present invention provides safe and easily controllabledelivery systems, based on particular biodegradable hydrogels, whichincrease the applicability of protein drugs for the treatment of variousdiseases. The risks associated with these drugs, such as bursts in therelease profile, and the inconvenience for the patient are reduced,while the therapeutic efficacy of drug treatments using the hydrogels ofthe present invention is increased.

[0022] More in detail, the present invention relates to a biodegradablehydrogel comprising bonds which are hydrolysable under physiologicalconditions. The hydrogels of the present invention containhydrolytically labile spacers, which spacers are broken in human oranimal bodies. As roughly indicated herein-above, a hydrogel is definedas a water-swollen, three-dimensional network of crosslinked hydrophilicmacromolecules. More in detail, the hydrogels of the invention consistof two interpenetrating networks interconnected to one another throughhydrolysable spacers. Hydrogels generally contain from 20 to more than99 wt. % water.

[0023] Further, the invention relates to a method for the preparation ofa hydrogel, wherein macromolecules, e.g. polymers, which contain bondswhich are hydrolysable under physiological conditions, are crosslinkedin an aqueous solution.

[0024] In accordance with the present invention, all types ofbiodegradable hydrogels can be used, provided that hydrolytically labilespacers can be introduced in these structures. When brought into thesystem of an animal or human body, the hydrogel structure is more orless gradually broken down. The degradation products do not need to beremoved after therapy; they can simply be metabolized and/or excreted bythe body.

[0025] Preferably, the hydrogels are based on water soluble polymerswhich contain at least a number of side groups having the capability toform linkers to other polymers, e.g. dextran or derivatized dextrans,while starches and starch derivatives, cellulose derivatives such ashydroxyethyl and hydroxypropyl cellulose, polyvinylpyrrolidone,proteins, polyamino acids, polyvinylalcohol, polyacrylates,polymethacrylates, polyethylene glycols that contain a number of monomerunits capable of forming side chains, and so on, can also be used, aswell as copolymers thereof.

[0026] The hydrogels of the invention are suitably based on a polymercrosslinked with methacrylate units. Other crosslinking units areacrylate units, vinyl ethers and vinyl esters, as well as other unitsknown for this purpose by the person skilled in the art.

[0027] Generally, the water-swellable polymers used in the presentinvention are made hydrolysable by introducing at least onehydrolytically labile unit in the spacers between the main chains ofwater soluble polymers and the second polymer chain formed by thecrosslinkable units as described in the preceding paragraph. It is alsopossible to use polymer chains comprising hydrolytically labile monomerunits in the main chain. However, not within the scope of the presentinvention are hydrogels which only contain polymer chains that areinterconnected head-to-tail only by hydrolyzable spacer groups. Contraryto hydrogels according to the invention, in hydrogels based on swellablepolymers which are only substituted with labile spacers head-to-tail,the degree of crosslinking is directly correlated to the water contentof the gel. The polymer system the hydrogel of the invention, makes itpossible to control the release of compund by adjusting the watercontent and/or the degree of crosslinking independently.

[0028] In a preferred embodiment, the hydrogels of the present inventioncomprise as hydrolytically labile units hydrolysable lactate and/orcarbonate ester bonds. These bonds can be brought into the hydrogel byintroducing e.g. (poly)glycolic acid and/or (poly)lactic acid residuesbetween the main chain of the polymer and the crosslinkable groups ofsaid polymer. With the term (poly)glycolic acid, glycolic acid as wellas di- and oligomers thereof are meant. With the term (poly)lactic acid,lactic acid as well as di- and oligomers thereof are meant. Otherpossibilities for hydrolytically labile units are based on unitsintroducing carboxylic esters, urethanes, anhydrides, (hemi)acetals,amides and peptide bonds.

[0029] In the most preferred embodiment of the present invention,(poly)glycolic acid and/or (poly)lactic acid spacers are introducedbetween polymerizable methacrylate groups and dextran. When a hydrogelformed of this material is introduced in a physiological environment,the hydrogel becomes biodegradable resulting in dextran,polyhydroxyethylmethacrylate (PHEMA), lactic acid and/or glycolic acidas degradation products. These degradation products are allbiocompatible. It is noted that lactic acid and glycolic acid areendogenous compounds. Dextran is a non-toxic polymer, which is used formany years as plasma expander and is cleared by the kidneys depending onits molecular weight. PHEMA is a well-known biocompatible polymer, whichis probably cleared via the kidneys, as well.

[0030] Hydrogels of the present invention can suitably be prepared byfirst synthesizing spacers which contain at least one crosslinkablegroup and at least one hydrolytically labile group; coupling suchspacers to a water-soluble polymer, and crosslinking the polymersobtained, preferably in the presence of the compound to be released.

[0031] Preferred spacers within the present invention comprise ahydroxyethyl methacrylate (HEMA) group, coupled to one or more lactideand/or glycolide units, as exemplified in steps a and b of scheme 1.More in detail, HEMA-terminated polylactic acid pre-polymers can besynthesized by solution polymerization of lactide in toluene using HEMAas initiator and aluminium alkoxide as catalyst or by a bulkpolymerization of lactide using HEMA as initiator and stannous octoateas catalyst. Dependent on the ratio HEMA/lactic acid pre-polymers whichdiffer in molecular weight of the lactic acid block can be synthesized.HEMA terminated copolymers of glycolic acid and of glycolicacid-co-lactic acid can be synthesized in analogy. The pre-polymers canbe characterized by known techniques, e.g. NMR and IR spectroscopy,differential scanning calorimetry (DSC), and gel permeationchromatography (GPC). In order to couple the HEMA terminated polylacticand/or glycolic acid pre-polymers to dextran, the terminal hydroxylgroup has to be activated. Preferably, the binding of HEMA to dextran iseffected by carbonyl-di-imidazole (CDI) as coupling agent. However, alsoother activation methods can be used. For example, reaction of thehydroxyl function of the HEMA lactic acid pre-polymer with succinicanhydride, followed by activation of the formed carboxylic group usingestablished methods (e.g. dicyclohexylcarbodiimide (DCC) activation).The latter method yields dextran derivatives in which onlyhydrolytically instable ester bonds are present, which provide differentdegradation characteristics compared to the dextran derivativessynthesized with the CDI-method, in which both ester bonds and carbonatebonds are present.

[0032] The activated HEMA terminated polylactic and/or glycolic acidpre-polymers are subsequently coupled to dextran in a suitable aproticsolvent (such as DMSO), in the presence of a catalyst, e.g.,N,N-dimethylamino-pyridine (DMAP). The degree of substitution (i.e.number of moles of methacrylate groups containing prepolymer per 100moles glucose units of dextran) can be tailored by the ratio of HEMAcontaining prepolymer versus dextran in the reaction mixture.

[0033] From these substituted polymers, hydrogels are prepared, e.g., bya radical polymerization of an aqueous solution of HEMA-pre-polymerfunctionalized dextran using a known initiator system of a tertiaryamine and a persulphate (see e.g. the above cited articles of VanDijk-Wolthuis et al. in Macromolecules 28, and Hennink et al. in J.Contr. Rel. 39). It is also possible to use polymerization by gammairradiation, with the advantage that no initiator and/or catalystresidues have to be extracted from the hydrogel.

[0034] The hydrogel system of the present invention can easily betailored with respect to protein drug release kinetics, whichtremendously expand the applicability of protein drugs. Especially, inthe case of biological response modifiers with a narrow therapeuticwindow, which are useful in the treatment of various diseases whereinthe immune system is involved, this is very important.

[0035] An increasing degree of substitution (DS; amount of hydrolysablespacer containing crosslinkable branches per 100 main water-solublepolymer residues; determinable by ¹H-NMR) yields a more crosslinkednetwork. This results in a slower swelling rate and an increasingdissolution time of the gel.

[0036] The hydrogels of the present invention can be prepared in such away that dissolution times from less than 1 day upto about 3 months andlonger can be obtained. This can for instance be effected by varying theinitial water content in the aqueous polymer solution to be crosslinkedand the DS. Gels with a high initial water content, such as watercontents higher than 85 wt. % predominantly contain intramolecularcrosslinks, while lower initial water contents give more intermolecularcrosslinking. Gels with less intermolecular crosslinks dissolve fasterat the same DS.

[0037] Drugs can be loaded into hydrogels either by equilibration in adrug-containing solution (see e.g. Kim et al. in Pharm. Res. 9(3) (1992)283-290) or by incorporation of the drug during the preparation ofhydrogel (see e.g. Heller et al. in Biomaterials 4 (1983) 262-266).

[0038] Loading by equilibration, however, normally leads to a rather lowdrug content in the delivery system. This is especially the case, whenthe drug is a macromolecular compound. Unless the pore size of thehydrogel is rather large, the macromolecules will only adhere to theouter surface, which may lead to a burst release.

[0039] Therefore, preferably, the drug is loaded during polymerizationor crosslinking.

[0040] Since microsphere suspensions are easy to prepare and are easilyused for injection, the hydrogels generally will be applied asmicrospheres of varying sizes. Microspheres can be prepared bydissolving HEMA-derivatized dextran in water after which this solutionis added to an oil phase (e.g. silicone oil) yielding a water-in-oilemulsion. After addition of a suitable initiator system, themethacrylate groups polymerize, yielding stable micropheres.

[0041] The drugs are released from the biodegradable hydrogels of thepresent invention during hydrolysis of the hydrogel, although at leastto some extent diffusion of proteins from the hydrogel takes place. Infact, the hydrolysis behaviour of the hydrogel and the time during whichcompounds present in the hydrogel system are released can be adjusted toone another so that the release can take place at any level betweenfirst order (no degradation of the hydrogel) and zero order release(fully degradation controlled) (see in this respect Example 6,herein-below). This provides evident advantage over hydrogel systemsthat are not hydrolysable at physiological conditions, but which arerather stable, such as the known dextran-GMA hydrogel system, or oversystems wherein the polymers in the hydrogel are elongated in onedimension only.

[0042] Protein drugs are released from the rather stable hydrogelsfollowing first order kinetics— protein release proportional to thesquare root of time— which is common for monolithic delivery systems.The hydrogels of the present invention, however, show a more zero orderrelease behaviour— protein release proportional to time. When thehydrogel degrades during the release process, the diffusion coefficientof the protein drug, present in the hydrogel, increases. This leads to amore constant release in time.

[0043] The hydrogel system of the present invention offers—as said—thepossibility to tailor the release profiles of encapsulated proteindrugs. More in detail, the degradation rate of the hydrogel can beadjusted by varying the water content of the hydrogel, the degree ofsubstitution, the number and length of hydrolysable groups in thespacers, and the choice of hydrolysable spacers.

[0044] It has been found that spacers based on glycolic acid have ahigher hydrolytical sensitivity than spacers based on lactic acid. Ifglycolic acid based spacers are present an accelerated degradation rateof the hydrogel is observed as compared with lactic acid based spacers.

[0045] The effect of the water content of the hydrogel and the degree ofsubstitution of the hydrogel polymers is elaborated in the examplesfollowing below.

[0046] Further, the rate of release depends on the size of the hydrogelparticles. This size can be adjusted by varying the stirring speed,viscosity of the external phase etc. during the preparation ofmicrospheres.

[0047] The rate of release does not depend on the length of the watersoluble polymers, at least not to a high extent. This is contrary tohydrogel systems wherein the hydrolysable groups are present in the mainchains of the polymer only (one dimensionally elongated polymers).

[0048] As indicated herein-above, the releasable compound can be aprotein drug. However, it is also possible to encapsulate pharmaconcontaining nanoparticles, e.g. liposomes, iscoms, polylactic acidparticles, polycaprolacton particles and gene delivery systems known tothe person skilled in the art. The encapsulation of these nanoparticleshas the advantage of preventing the occurence of a too fast release ofthe encapsulated compound, or, said in other words, burst-effects can beavoided in a more secure way.

[0049] An example of a loaded hydrogel system within the presentinvention is a hydrogel containing the cytokine interleukin-2 (IL-2).IL-2 is a protein drug which can e.g. be used in the treatment ofparticular types of cancer.

[0050] For IL-2 to be therapeutically effective in cancer treatment,prolonged presence of IL-2 at the site of tumor growth is required. Thiscan be achieved either by administering high doses of IL-2 intravenouslythrough frequent bolus injections (see e.g. Rosenberg et al. JAMA 271,(1994) 907-913), by prolonged continuous infusion (see e.g. West et al.,N. Engl. J. Med. 316, (1987), 898-905, or by frequently administeringlow doses of IL-2 intra- or peri-tumorally (see e.g. Den Otter et al.;Anticancer Res. 13, (1993), 2453-2455).

[0051] A major disadvantage of the intravenous route is that forobtaining sufficiently high levels of IL-2 at the site of tumor growth,such high doses of IL-2 have to be administered intravenously, that itbecomes severely toxic. In contrast, the intra- or peri-tumoralapproach, as developed by Den Otter et al., has proven to be verysuccessful and virtually non-toxic in various transplanted andspontaneous tumors.

[0052] A serious problem for application of this form of therapy inhuman cancer patients, however, is that IL-2 has to be injected intra-or peri-tumorally 5 to 10 times within 1 to 2 weeks. For many types ofcancer this is not-acceptable burden for the patient, like in cases oflung carcinoma, bladder cancer, and gastric cancer. In first attempts totranslate the very effective local, low-dose IL-2 treatment to the humancancer clinic, these logistic problems of IL-2 delivery were already runinto. The slow-release delivery system of the present invention makesthe use of local IL-2 immunotherapies possible.

[0053] For the in vivo application hydrogel suspensions (microspheres)will normally contain up to 10⁵ I.U. of IL-2 in 0.5 ml, which arereleased over a period of 5 days (i.e. 2×10⁴ I.U. of IL-2 released perday). The amount of protein released can be determined with sensitivequantitative detection methods (HPLC, ELISA assays). To investigatewhether the released IL-2 is still biologically active and to whatextent (i.e. what is the effect of these chemical procedures on thespecific activity of IL-2), proliferation assays using the IL-2dependent CTLL cell line can be performed.

[0054] The present invention will now be explained in more detail, whilereferring to the following, non-limiting examples.

EXAMPLES

[0055] To obtain a dextran hydrogel, first a polymerizable methacrylategroup was introduced in dextran. For all reactions described below,dextran from Fluka (T40, M_(n)=15.000, M_(w)=32.500 g/mol) was used.Four different dextran derivatives were synthesized (Examples 1-4) inwhich the methacrylate ester was coupled via a spacer to dextran. Thespacer contains different hydrolyzable bonds (carbonate and/orcarboxylic ester). The degradation behaviour of gels prepared from thesederivatives was compared with dextran gels derived fromglycidylmethacrylate (dex-GMA). In this reference compound, themethacrylate ester is directly coupled to a hydroxyl group of dextran.This reference gel degrades extremely slowly.

EXAMPLE 1 Synthesis of dex-HEMA

[0056] Dextran derivatized with hydroxyethyl methacrylate (dex-HEMA) wassynthesized by coupling carbonyldiimidazole (CDI) activated HEMA(HEMA-CI) to dextran.

[0057] CDI (1.62 g; 10 mmol) was dissolved in about 10 ml anhydroustetrahydrofuran (THF). This solution was added to a solution of HEMA(1.30 g; 10 mmol) in 5 ml anhydrous THF. The reaction mixture wasstirred for 16 hours at room temperature. After evaporation of thesolvent a slightly yellow liquid was obtained (yield 2.93 g). The crudeproduct was dissolved in ethylacetate, extracted with water to removeimidazole and unreacted HEMA and dried on anhydrous MgSO₄. Afterfiltration, the solvent was evaporated and almost pure hydroxyethylmethacrylate activated with CDI (HEMA-CI) was obtained. The structure ofthis product was confirmed by NMR and IR spectroscopy.

[0058] Varying amounts of HEMA-CI (0.73, 1.49, or 2.19 g; 96% pure) wereadded to a solution of dextran (10 g, 62 mmole glucose units) andN,N-dimethylaminopyridine (DMAP; 2 g, 16.4 mmol) in anhydrousdimethylsulfoxide (DMSO; 90 ml). These reaction mixtures were stirredfor 4 days at room temperature after which the reaction was terminatedby the addition of about 2 ml of concentrated HCl. The reaction mixturewas dialyzed against water for 3 days at 4° C. Dex-HEMA was isolated bylyophilization yielding a white fluffy material (yield>90%). The degreeof HEMA substitution was determined by NMR spectroscopy, and amounted 4,9, and 13, respectively.

EXAMPLE 2 Synthesis of dex-SA-HEMA

[0059] Dextran derivatized with the hemi-ester of succinic acid (SA) andHEMA (dex-SA-HEMA) was synthesized as follows.

[0060] SA (2.00 g, 20 mmol), HEMA (2,6 g, 20 mmol), triethylamine (TEA;0.28 ml, 2 mmol) and hydrochinon (inhibitor, ±10 mg) were dissolved inabout 30 ml anhydrous THF. The reaction mixture was stirred for 2 daysat 45° C., after which the solvent was evaporated. A yellow liquid wasobtained (yield 4.88 g). The structure of HEMA-SA was confirmed by NMRand IR spectroscopy.

[0061] HEMA-SA (0.99 g (94% pure), 4 mmol) and dicyclohexylcarbodiimide(DCC; 0.83 g, 4 mmol) were dissolved in about 20 ml anhydrous DMSO.After 15 minutes a precipitate was formed (dicyclohexylureum; DCU) andthis mixture was added to a solution of dextran (2.57 g, 16 mmol glucoseunits) and TEA (0.56 ml, 4 mmol) in anhydrous DMSO (20 ml). Theresulting mixture was stirred for 3 days at room temperature, afterwhich 3 drops of concentrated HCl were added to terminate the reaction.Next, the reaction mixture was filtered to remove DCU and dialyzed for 3days at 4° C. After lyophilization, a white fluffy product was obtained(yield 2.78 g). The degree of substitution was established by NMRspectroscopy and amounted to 3.

EXAMPLE 3 Synthesis of dex-lactate-HEMA

[0062] Dextran derivatized with HEMA-oligolactide was synthesized asfollows as illustrated in scheme 1. Three steps can be distinguished.

[0063] a. coupling of lactate to HEMA yielding HEMA-lactate;

[0064] b. activation of, HEMA-lactate using CDI yielding HEMA-lactate-CI

[0065] c. coupling of HEMA-lactate-CI to dextran.

[0066] A mixture of L-lactide (4.32 9; 30 mmol) and HEMA (3.90 g; 30mmol) was heated to 110° C. Thereafter, a catalytic amount of stannousoctoate (SnOct₂; 121.5 mg, 0.3 mmol) dissolved in about 0.5 ml toluenewas added. The resulting mixture was stirred for 1 hour. After coolingdown to room temperature, the mixture was dissolved in THF (20 ml). Thissolution was dropped in water (180 ml) and the formed precipitate wasisolated by centrifugation. The pellet was taken up in ethyl acetate (40ml), centrifugated to remove solid material, dried (MgSO₄) and filtered.The solvent was evaporated yielding a viscous oil (3.74 g, 45%). Theproduct (HEMA-lactate) was characterized by NMR and IR spectroscopy.

[0067] HEMA-lactate (3.74 g, 10.8 mmol) was added to a solution of CDI(1.76 g, 10.8 mmol) in THF and stirred for 16 hours at room temperature.The solvent was evaporated under reduced pressure yielding a viscousoil. The product containing HEMA-lactate-CI and imidazole as majorcompounds (NMR analysis) was used without further purification.

[0068] To a solution of dextran (10 g, 62 mmol glucose units) and DMAP(2.0 g, 10.6 mmol) a varying amount of HEMA-lactate-CI was added (1.61,3.23 or 4.84 g respectively, 80% pure). These mixtures were stirred for4 days at room temperature after which the reaction was terminated bythe addition of about 2 ml of concentrated HCl. The solutions weredialyzed against water for 2 days. After lyophilization, white fluffyproducts were obtained (yield around 85%). The degree of substitution(as determined by NMR spectroscopy) amounted to 3, 6 and 10 for thethree products, respectively.

[0069] Using similar procedures, the length of the lactate spacer can bevaried by changing the molar ratio of HEMA and lactide in the firstreaction.

EXAMPLE 4 Synthesis of dex-glycolide-HEMA

[0070] Dex-glycolide-HEMA was synthesized according to the sameprocedure as described in EXAMPLE 3 replacing lactide by glycolide.

Reference Example 1 Dex-GMA

[0071] Dex-GMA was synthesized as described by Van Dijk-Wolthuis et al.,Macromolecules 28, (1995), 6317-6322. An reinvestigation of the obtaineddextran derivative revealed that the methacrylic ester group is directlycoupled to one of the hydroxylic group of dextran, meaning that theglyceryl spacer is not present.

EXAMPLE 5 Preparation of Dextran Hydrogels

[0072] Hydrogels were obtained by a free radical polymerization ofaqueous solutions of methacrylated dextran (prepared according toEXAMPLES 1-4 and Ref. Example 1). Methacrylated dextran (100 mg) wasdissolved in 760 μl PBS buffer (10 mM phosphate, 0.9% NaCl, 0.02% NaN3,pH 7.2). To this solution 90 μl of potassium peroxydisulfate (KPS; 50mg/ml) in the same buffer was added per gram solution and mixed well.Subsequently, N,N,N′N′-tetramethylethylenediamine (TEMED; 50 μl; 20%(v/v) in water, pH adjusted to 7) was added and the resulting solutionwas quickly transferred into an Eppendorf tube and polymerized for 1hour at room temperature yielding a hydrogel material with an initialwater content of about 90% (w/w) after polymerization.

[0073] The gels were removed from the tubes, accurately weighed andincubated in PBS at 370° C. Periodically, the weight of the gels wasdetermined and used to calculate the degree of swelling (=W_(t)/W_(o),in which W_(t) is the weight of the gel at time t and W_(o) is theinitial weight of the gel). The hydrogel degradation (dissolution) timeis defined as the time at which the swelling ratio=0 (or W_(t)=0).

[0074]FIG. 1 shows the swelling behaviour of three different dextranhydrogels (initial water content 90%). The dex-GMA (DS=4) reached anequilibrium swelling within 3 days; thereafter the weight of the gelsremained constant indicating that no significant hydrolysis ofmethacrylate esters occurred. Dex-HEMA and dex-lactate-HEMA showed aprogressive swelling in time until these gels dissolved completely. Thisdemonstrates that in these hydrogel systems hydrolysis of carbonateesters (dex-(lactate)HEMA) and/or lactate esters (dex-lactateHEMA)occurred.

[0075]FIG. 2 shows the swelling behaviour of dex-lactateHEMA hydrogels(initial water content 92%) and varying degree of substitution. As canbe seen an increasing DS resulted in an increasing dissolution time.

[0076]FIG. 3 shows the swelling behaviour of dex-lactateHEMA hydrogelswith a fixed DS (6) and varying initial water content. It appears thatthe dissolution time increases with an increasing initial water content.

[0077]FIG. 4 shows the swelling behaviour of dex-SAHEMA hydrogel (DS 3)and a varying initial water content.

[0078] The next tables give an overview of the dissolution times ofdifferent dextran hydrogels dissolution time (days) of dex-lactate-HEMAgels initial water content DS 10 DS 6 DS 2.5 90% 8-13 4 1-2 80% 16  7-183-9 70% 22 10-19  6-10 dissolution time (days) of dex-HEMA gels initialwater content DS 8 DS 17 DS 20 92% 25 42 55 80% 53 100 ND 67% 70 >100 NDdissolution time (days) of dex-HEMA-SA gels initial water content DS 380% 15 74% 23-28 60% 44

[0079] The dissolution time of dextran hydrogels (initial water content92%) obtained by polymerization of dextran derivatized with GMA (DS 4)had a dissolution time of about 100 days (pH 7.2, 37° C.). Gels obtainedby polymerization of dextran derivatized with GMA (initial water content80%, DS 11) did not show any signs of degradation (increased swelling)during 70 days, even at extreme conditions (37° C., pH 12 and pH 1).

EXAMPLE 6 Protein Release from Degrading Dextran Hydrogels

[0080] The release from non-degrading dex-GMA hydrogels has been studiedextensively. It appeared that when the protein diameter was smaller thanthe hydrogel mesh size, the release of the protein could be effectivelydescribed by the free volume theory. In this case the cumulative amountof protein released was proportional to the square root of time (W. E.Hennink, Controlled release of proteins from dextran hydrogels, Journalof Controlled Release 39, (1996), 47-55).

[0081]FIG. 5 shows the release of a model protein (IgG) from degradingdextran hydrogels (dex-lactate-HEMA, DS 2.5). The release of IgG fromthese degrading gels is zero order (cumulative release proportional totime). This is in contrast with the release of proteins fromnon-degrading dextran hydrogels, where a first order release has beenobserved.

1. Biodegradable hydrogel comprising bonds which are hydrolysable underphysiological conditions.
 2. Hydrogel according to claim 1, consistingof two interpenetrating polymer networks interconnected to one anotherthrough hydrolysable spacers.
 3. Hydrogel according to claim 1 or 2,being based on crosslinked dextran or a crosslinked derivatized dextran.4. Hydrogel of any one of the preceding claims, based on a polymercrosslinked with methacrylate units.
 5. Hydrogel of any one of thepreceding claims, comprising hydrolysable lactate and/or carbonate esterand/or succinic acid and/or ethylene glycol bonds.
 6. Method for thepreparation of a hydrogel, wherein macromolecules, e.g. polymers, whichcontain bonds which are hydrolysable under physiological conditions, arecrosslinked in an aqueous solution.
 7. The method of claim 6, whereinthe hydrolysable bonds are derived from lactic acid and/or glycolic acidand/or succinic acid and/or ethylene glycol.
 8. The method of claim 6 or7, wherein a protein drug is present during the crosslinking step.